This invention relates to the art of diagnostic imaging. It finds particular application in conjunction with dual energy, scanned detector array medical diagnostic digital radiography, and will be described with particular reference thereto. However, it is to be appreciated that the invention may find further application in conjunction with other imaging techniques such as single or multiple energy digital radiography, computed tomography, and the like.
In prior art digital radiography, each image is commonly a rectangular array of pixels. The pixel value, which is usually displayed as gray scale, is indicative of the degree of radiation attenuation or transmissivity through a corresponding path of the region of interest.
By conducting the same examination with two different energy levels, material specific images, called basis images, can be derived, in addition to a normal image which a linear combination of high and low energy images. With human patients, soft tissue and calcium specific images are most commonly generated.
It is to be appreciated that the accuracy of the basis images is dependent upon the degree to which they are determined from high and low energy radiation beams that have passed through the same tissue. When the high and low energy data has been collected serially, i.e. high energy data was collected followed by low energy data, registration of the two images is difficult. Patient motion between the images complicates the registration procedure and, in some cases, makes it impossible. Compensation can sometimes be made for simple, lateral patient movement by registering portions of the image data separately. A detector is defined as a phosphor screen, photodiode pair. Movement of the patient toward or away from the detector causes magnification errors.
To avoid patient movement problems, the high and low energy data can be collected concurrently. However, difficulty still arises in aligning the high and low energy detectors precisely. Sometimes, high and low energy detectors are separated by a filter such that lower energy radiation causes light emission primarily from the front phosphor and high energy radiation causes light emission from the rear phosphor. A first array of photodiodes or other photodetectors are disposed along the front phosphor screen and a second array of photodiodes or other radiation detectors are mounted to the rear phosphor screen. Commonly, photodiode chips or modules are manufactured with a plurality of photodiodes in a linear array, e.g. 32 photodiodes. A plurality of the arrays or modules are abutted and aligned until the entire examination region is spanned.
The phosphor screens are commonly arced such that all points along the screen are equidistant from the radiation source. Aligning the front and back photodiode arrays is rendered more difficult by the curvature of the phosphor screens. The front and back detectors tend to magnify the spatial position to different degrees. In some scanners, the thickness of the phosphor screens and the filter create a parallax problem in the alignment of the front and back diode arrays.
Mechanical alignment techniques provide relatively good alignment between the corresponding high and low energy photodiode arrays. Basis images are derived by a non-linear function of the high and low image values. Therefore misalignments of even a small fraction cause significant errors at a material boundary. The sign of the error depends on the direction of misalignment and direction of a change in material thickness. At abrupt material changes, both positive and negative errors occur in close proximity. These artifacts are commonly called phase change errors. The non-linear basis transformation that separates high and low energy data into material specific images exaggerates the phase change errors and the resultant artifacts in the material specific images. Because the photodiodes are commonly very small, e.g. on the order of one millimeter, even misalignments as little as a tenth of a millimeter or less cause significant errors in the basis images.
Analogous registration problems occur when a multilinear array of detectors is scanned. A multilinear detector array is made up of a plurality of multilinear chips abutted together forming an array of detector elements. Each chip is a 20 photodiode array composed of m columns x n rows of active elements which are integrated over time to produce signals effectively equivalent to a linear array of n elements. The detectors in each row sequentially shift or move into alignment with each pixel location. Each detector is sampled when in alignment with the pixel. The sampled data is summed with data from other detectors in the row sampled when each was in alignment with the same pixel. An example of a multilinear detector array utilizes the shift and add sequence can be found in Kruger U.S. Pat. No. Re. 32,164 which is incorporated herein by reference. It is to be appreciated that if the detectors in the row are not perfectly aligned, the data sampled from each will not represent radiation that has traversed the same path through the patient. Any misalignment of the sequential sampled detectors again causes image degradation.
In a single linear array detector system, the detectors are commonly wider than one pixel. Pre-patient collimation assures that the effective detector width is constant along the array. In one prior art alignment technique, two sequentially collected images are brought into alignment using translation and rotation. Patient movement may require different areas of the images to have different translational or rotational differences.
In accordance with the present invention, a new and improved electronic data processing technique is provided for correcting basis images for high and low energy photodiode pair misalignment.